Non-interferometric photoacoustic remote sensing (ni-pars)

ABSTRACT

A photoacoustic remote sensing system (NI-PARS) for imaging a subsurface structure in a sample, has an excitation beam configured to generate ultrasonic signals in the sample at an excitation location; an interrogation beam incident on the sample at the excitation location, a portion of the interrogation beam returning from the sample that is indicative of the generated ultrasonic signals; an optical system that focuses at least one of the excitation beam and the interrogation beam with a focal point that is below the surface of the sample; and a detector that detects the returning portion of the interrogation beam.

CROSS-REFERENCE TO RELATED APPLICATION(S

This Patent application is a continuation of U.S. Pat. Application No.17,091,856, filed on Nov. 6, 2020, which is a continuation of U.S. Pat.Application No. 16/402,972, filed on May 3, 2019, which is acontinuation of U.S. Pat. Application No. 15/418,447, filed on Jan. 27,2017, now U.S. Pat. No. 10,327,646, which claims priority to U.S.Provisional Application No. 62/290,275, filed Feb. 2, 2016, theentireties of which are incorporated herein by reference.

FIELD

This relates to the field of biomedical optics imaging and, inparticular, to a laser and ultrasound-based method and system for invivo or ex vivo, non-contact imaging of biological tissue.

BACKGROUND

Photoacoustic imaging is an emerging hybrid imaging technology providingoptical contrast with high spatial resolution. Nanosecond or picosecondlaser pulses fired into tissue launch thermo-elastic-induced acousticwaves which are detected and reconstructed to form high-resolutionimages. Photoacoustic imaging has been developed into multipleembodiments, including photoacoustic tomography (PAT), photoacousticmicroscopy (PAM), optical-resolution photoacoustic microscopy (OR-PAM),and array-based PA imaging (array-PAI). In photoacoustic tomography(PAT) signals are collected from multiple transducer locations andreconstructed to form a tomographic image in a way similar to X-ray CT.In PAM, typically, a single element focused high-frequency ultrasoundtransducer is used to collect photoacoustic signals. A photoacousticsignal as a function of time (depth) is recorded for each position in amechanically scanned trajectory to form a 3-D photoacoustic image. Themaximum amplitude as a function of depth can be determined at each x-yscan position to form a maximum amplitude projection (MAP) C-scan image.Photoacoustic microscopy has shown significant potential for imagingvascular structures from macro-vessels all the way down tomicro-vessels. It has also shown great promise for functional andmolecular imaging, including imaging of nanoparticle contrast agents andimaging of gene expression. Multi-wavelength photoacoustic imaging hasbeen used for imaging of blood oxygen saturation, by using known oxy-and deoxy-hemoglobin molar extinction spectra.

In traditional photoacoustic imaging, spatial resolution is due toultrasonic focusing and can provide a depth-to-resolution ratio greaterthan 100. In OR-PAM, penetration depth is limited to ∼1 mm in tissue(due to fundamental limitations of light transport) but resolution ismicron-scale due to optical focusing. OR-PAM can provide micron-scaleimages of optical absorption in reflection-mode, in vivo, something thatno other technique can provide. OR-PAM is capable of imaging bloodvessels down to capillary size noninvasively. Capillaries are thesmallest vessels in the body and so much crucial biology occurs at thislevel, including oxygen and nutrient transport. Much can go wrong at thecapillary level too. In cancers, cells have an insatiable appetite foroxygen and nutrients to support their uncontrolled growth. They invoke arange of signaling pathways to spawn new vessels in a process known asangiogenesis and these vessels typically form abnormally. Tumors areoften highly heterogeneous and have regions of hypoxia. Photoacousticimaging has demonstrated the ability to image blood oxygen saturation(SO2) and tumor hypoxia in vivo.

In most photoacoustic and ultrasound imaging systems, piezoelectrictransducers have been employed, in which an ultrasound coupling mediumsuch as water or ultrasound gel is required. However for many clinicalapplications such as wound healing, burn diagnostics, surgery, and manyendoscopic procedures physical contact, coupling, or immersion isundesirable or impractical.

The detection of ultrasound in photoacoustic imaging has, untilrecently, relied on ultrasonic transducers in contact with thebiological tissue or an ultrasonic coupling agent both of which havemajor drawbacks as described above. Some detection strategies to solvingthe non-contact optical interferometric sensing problems associated withphotoacoustic imaging have been reported.

Optical means of detecting ultrasound and photoacoustic signals havebeen investigated over a number of years; however, to date no techniquehas demonstrated practical non-contact in vivo microscopy in reflectionmode with confocal resolution and optical absorption as the contrastmechanism.

Most previous approaches detected surface oscillations withinterferometric methods. Others used interferometry to observephotoacoustic stresses, including optical coherence tomography (OCT)methods. These methods offer potential sensitivity to the scatteredprobe beam phase modulations associated with motion of scatterers,subsurface and surface oscillations, as well as unwanted vibrations.They are also sensitive to complex amplitude reflectivity modulations.

One example of a low-coherence interferometry method for sensingphotoacoustic signals was proposed in U.S. pregrant publication no.2014/0185055 to be combined with an optical coherence tomography (OCT)system, resulting in 30 µm lateral resolution.

Another prior art system is described in U.S. pregrant publication no.2012/0200845 entitled “Biological Tissue Inspection Method and System”,which describes a noncontact photoacoustic imaging system for in vivo orex vivo, non-contact imaging of biological tissue without the need for acoupling agent.

Other systems use a fiber based interferometer with opticalamplification to detect photoacoustic signals and form photoacousticimages of phantoms with acoustic (not optical) resolution. However thesesystems suffer from a poor signal-to-noise ratio, other contact-basedphotoacoustic systems offer significantly improved detectioncapabilities, in vivo imaging was not demonstrated, andoptical-resolution excitation was not demonstrated.

Industrial laser ultrasonics has used interferometry to detect acousticsignatures due to optical excitation of inanimate objects fornon-destructive testing. This approach has been adapted to detectultrasound ex vivo in chicken breast and calf brain specimens, however,optical-resolution focusing of the excitation light was not examined.

Laser Doppler vibrometry has been a powerful non-contact vibrationsensing methodology, however, weak signal-to-noise and poor imagequality have proven to be a limitation when sensing deep-tissue signalsfrom broad-beam photoacoustic excitation.

Similarly, Mach Zehnder interferometry and two-wave mixinginterferometry have been used previously for sensing photoacousticsignals. However many such techniques still require direct contact orfluid coupling; have not offered in vivo studies or optical resolutionfor phantom studies.

The non-interferometric photoacoustic remote sensing (NI-PARS) isfundamentally different from other approaches for detectionultrasound/photoacoustic signals. The system takes advantage of a pulsedexcitation beam co-focused and co-scanned with an interrogation beam.The detection mechanism is based on a non-interferometric sensing.Rather than detecting surface oscillations, pressure-inducedrefractive-index modulation resulting from initial pressure fronts canbe sampled right at their subsurface origin where acoustic pressures arelarge. The non-interferometric nature of detection along with theshort-coherence lengths of the interrogation laser preclude detection ofsurface- and sub-surface oscillations to provide only the initialpressure signals.

SUMMARY

According to an aspect, there is provided a non-interferometricphotoacoustic remote sensing system (NI-PARS) for imaging a subsurfacestructure in a sample, where the NI-PARS comprises an excitation beamconfigured to generate ultrasonic signals in the sample at an excitationlocation; an interrogation beam incident on the sample at the excitationlocation, a portion of the interrogation beam returning from the samplethat is indicative of the generated ultrasonic signals; an opticalsystem that focuses the excitation beam at a first focal point and theinterrogation beam at a second focal point, the first and second focalpoints being below the surface of the sample; and an optical detector todetect the returning portion of the interrogation beam.

According to another aspect, there is provided an endoscopic device thatuses a non-interferometric photoacoustic remote sensing system (NI-PARS)for imaging a subsurface structure in a sample, the endoscopic devicecomprising a fiber optic cable having an input end and a detection end;an excitation beam coupled to the input end of the fibre optic cable,wherein in use the excitation beam generates ultrasonic signals in thesample at an excitation location that is adjacent to the detection endof the fiber optic cable, the fiber optic cable focusing the excitationbeam at a first focal point that is below the surface of the sample; aninterrogation beam coupled to the input end of the fibre optic cable andincident on the excitation location, the fiber optic cable focusing theexcitation beam at a first focal point that is below the surface of thesample, and wherein a portion of the interrogation beam that isindicative of the generated ultrasonic signals is received by thedetection end of the fiber optic cable and travels to the input end; andan optical detector that receives the returning portion of theinterrogation beam at the input end of the fiber optic cable.

According to other aspects, either alone or in combination, asapplicable: the first and second focal points may be within 1 mm of thesurface of the sample; the first and second focal points may be greaterthan 1 µm below the surface of the sample; the focal point may be spacedbelow the surface of the sample at a depth that is greater than a focalzone of the respective at least one of the excitation beam and theinterrogation beam; the excitation beam and the interrogation beam havea lateral separation of less than 1 mm or less than 0.5 mm on thesample; the excitation beam may have a focal point that is laterallywithin the focal zone of the interrogation beam; the interrogation beammay have a focal point that is laterally within the focal zone of theexcitation beam; there may be a processor that calculates an image ofthe sample based on the returning portion of the interrogation beam; atleast one of the first focal point and the second focal point may have afocal diameter of less than 30 µm, 10 µm, or 1 µm; the excitation beammay be scanned through the sample while the interrogation beam isstationary; the interrogation beam may be scanned through the samplewhile the excitation beam is stationary; and each of the interrogationbeam and the excitation beam may be scanned through the sampleconcurrently.

The proposed NI-PARS approach intentionally eliminates phase-sensitivitydue to interferometric detection to exclusively monitor intensityreflectivity changes.

In one aspect, the approach generates high initial photoacousticpressures modify scattering properties, such as scattering cross sectionof individual particles or reflectivity from larger structures in thesample. This results in time-varying intensity reflectivity which doesnot require phase sensitive detection. Rejecting interferometric effectsleads to high signal to noise ratio detection.

To observe such reflection modulations, the intensity changes of a probebeam in response to a generated photoacoustic initial pressure aremeasured A non-interferometric approach with a low-coherence probe beamprecludes any phase-modulation sensitivity to enable detection ofintensity variations. The proposed approach transiently amplifiesexisting refractive index steps where absorption is present.

Other aspects will be apparent from the description and claims below.

BRIEF DESCRIPTION OF THE DRAWINGS

These and other features will become more apparent from the followingdescription in which reference is made to the appended drawings, thedrawings are for the purpose of illustration only and are not intendedto be in any way limiting, wherein:

FIG. 1 -FIG. 3 are the block diagram of non-interferometricphotoacoustic remote sensing (NI-PARS) microscopy systems

FIGS. 4 a-4 d are representative drawings of the overlap between theexciter and interrogator beams on a sample.

FIGS. 5 a - 5 c are block diagrams of examples of sensing systems in anendoscopy configuration.

FIG. 6 is a block diagram of a sensing system integrated with anotheroptical imaging system.

FIG. 7 a is a NI-PARS image of a network of carbon fibres.

FIG. 7 b is a graph of the FWHM obtained by fitting an individual carbonfiber signal amplitude to a Gaussian function.

FIG. 7 c is a graph of the resolution using a knife edge spreadfunction.

FIG. 8 in vivo images of CAM-membrane of 5-day chicken embryos.

FIG. 9 depicts an example of the frequency response of a NI-PARS system.

FIG. 10 depicts an example of the photoacoustic time domain signal of anindividual carbon fiber using a NI-PARS system.

FIG. 11 depict in vivo NI-PARS images of a mouse ear.

DESCRIPTION

Photoacoustic imaging is an emerging biomedical imaging modality thatuses laser light to excite tissues. Energy absorbed by chromophores orany other absorber is converted to acoustic waves due to thermo-elasticexpansion. These acoustic signals are detected and reconstructed to formimages with optical absorption contrast. Photoacoustic imaging (PA) hasbeen shown to provide exquisite images of microvessels and is capable ofimaging blood oxygen saturation, gene expression, and contrast agents,among other uses. In most PA and ultrasound imaging systemspiezoelectric transducers have been employed, in which an ultrasoundcoupling medium such as water or ultrasound gel is required. However formany clinical applications such as wound healing, burn diagnostics,surgery, and many endoscopic procedures physical contact, coupling, orimmersion is undesirable or impractical. The system described herein iscapable of in vivo optical-resolution photoacoustic microscopy usingnon-contact non-interferometric sensing without use of any ultrasoundmedium.

The system described herein, a non-interferometric photoacoustic remotesensing (NI-PARS) microscopy system, is based on the idea of focusingexcitation light to a near diffraction-limited spot and detectingphotoacoustic signals using a confocal interrogation beam co-focusedwith the excitation spot. While previous approaches for non-contactdetection of photoacoustic signals used interferometry detection as wellas a broad excitation beam with powerful lasers delivering mJ-J of pulseenergy over a broad area, the NI-PARS microscopy technique describedherein uses nJ-scale pulse energies focused to near diffraction-limitedspots. When focusing into tissue, the surface fluence can be maintainedbelow present ANSI limits for laser exposure but theballistically-focused light beneath the tissue can create fluencestransiently far above the ANSI limits (as is done in other microscopymethods). In NI-PARS, this means that very large local fluences ~J/cm²are created within a micron-scale spot, generating very large initialacoustic pressures. For example, at 532-nm excitation wavelength,imaging a capillary with 500 mJ/cm² local fluence would result in aninitial pressure on the order of 100 MPa locally. In NI-PARS approach,large optically-focused photoacoustic signals are detected as close tothe photoacoustic source as possible, which is done optically byco-focusing an interrogation beam with the excitation spot.

The major difference of our proposed work with previously publishedsystems is that a non-interferometric detection mechanism based onpressure-induced refractive-index modulation is used. Unlikeinterferometric methods, NI-PARS do not offer sensitivity to thescattered probe beam phase modulations associated with motion ofscatterers, subsurface and surface oscillations, as well as unwantedvibrations. The net interferometric signal may be a mixture of thesecomposite mechanisms and could lead to unwanted interference. Anon-interferometric approach with a low-coherence probe beam precludesany phase-modulation sensitivity to enable detection of intensityvariations. The proposed approach transiently amplifies existingrefractive index steps where absorption is present. This generates adetectable change in the reflection characteristic of a sample. In thecase of a surface much larger than the focal spot size of the detectionbeam this is a change in the intensity reflectivity of the surface. Inthe case of an object on the scale of, or smaller than, the detectionfocal spot size, this is a change in the scattering properties of theobject. This in turn effects the back reflected collected fraction, orin the case of a large collection of small objects will affect thescattering properties of the excited medium.

Since we do not have to perform depth scanning, NI-PARS can perform nearreal time using a high pulse repetition laser and fast scanning mirrors.However, most previous non-contact photoacoustic detection methods havenot shown real-time imaging capability and optical resolution was notdemonstrated. We optically focus a pulsed excitation laser intosuperficial tissues to generate high micro-scale initial pressures. Thenwe harvest these large optically-focused photoacoustic signals as closeto the photoacoustic source as possible. This is done by detectingphotoacoustic signals using a confocal interrogation beam co-focused andco-scanned with the excitation spot. Local initial pressures are verylarge when optical focusing and thermal confinement conditions areapplied. These large initial pressures can cause significant refractiveindex mismatch regions which are measured by the NI-PARS system aschanges in reflected light.

To the best of our knowledge this is the first report onultrasound/photoacoustic imaging detection mechanism based onpressure-induced refractive-index modulation as well as real-timenon-contact detection. Our approach is the only method we know of tointerrogate subsurface absorption with optical resolution using anon-contact system, aside from our previously disclosed interferometricPARS system, which however had a mixture of detection mechanisms, unlikethe current method which only senses initial pressure at subsurfacelocations.

The high sensitivity and the fine resolution of the proposed systemoffer performance comparable to other in vivo optical resolutionphotoacoustic microscopy systems but in a non-contact reflection modesuitable for many clinical and pre-clinical applications.

The general experimental setup of the non-interferometricoptical-resolution photoacoustic remote sensing microscopy system aredepicted through FIGS. 1-3 . Variations to the depicted system will beapparent to those skilled in the art. Referring to FIG. 1 , a blockdiagram of NI-PARS system 10, and in particular, a non-interferometricoptical-resolution photoacoustic remote sensing (NI-OR-PARS) microscopysystem, is shown. A multi-wavelength fiber excitation laser 12 is usedin multi focus form to generate photoacoustic signals. Excitation laser12 preferably operates in the visible spectrum, although the particularwavelength may be selected according to the requirements of theparticular application. The excitation beam 17 and interrogation beam 16pass through a lens system 42 to adjust their focus on the sample 18.The excitation beam 17 will be combined with interrogation beam 16 usinga beam combiner 30. The acoustic signatures are interrogated usingeither a short or long-coherence length probe beam 16 from a detectionlaser 14 that is co-focused and co-aligned with the excitation spots onsample 18. Interrogation/probe beam 16 passes through a polarizing beamsplitter 44 and quarter wave plate 56 to guide the reflected light fromsample 18 to the photodiode 46. The combined beam will be scanned byscanning unit 19. The scanning combined beams will pass through anobjective lens 58 and focused on the sample 18. The reflected beamreturns along the same path and is analyzed by detection unit 22. Unit22 consists of amplifier 48, fast data acquisition card 50 and computer52.

FIG. 2 shows the experimental setup on NI-PARS when the scanning unit 19is replaced by scanning unit 11 in order to scan the sample 19 relatedto the to the fixed combined beams 14.

FIG. 3 shows an experimental setup that excitation and interrogationbeams have separated path and are not combined. In this case theinterrogation beam will be focused using another objective lens 15 tothe sample 18.

In all the configurations, both beam can be scanned together. One beamcan be fixed while the other beam can be scanned. The sample 18 can bescanned while both beam are fixed. The sample 18 can be scanned whileboth beam are scanning. The sample 18 can be scanned while one beam isfixed and the other is scanning.

It will be apparent that other examples may be designed with differentcomponents to achieve similar results. For example, other examples couldinclude all-fiber architectures where circulators replace beam-splitterssimilar to optical-coherence tomography architectures. Otheralternatives may include various coherence length sources, use ofbalanced photodetectors, interrogation-beam modulation, incorporation ofoptical amplifiers in the return signal path, etc.

The NI-PARS system takes advantage of two focused laser beams on thesample which can simulate a confocal NI-PAM configuration.

The NI-PARS takes advantage of optical excitation and detection whichcan help dramatically reduce the footprint of the system. The absence ofa bulky ultrasound transducer makes this all optical system suitable forintegrating with other optical imaging systems. Unlike many previousnon-contact photoacoustic imaging systems, the NI-PARS system is capableof in vivo imaging. It relies on much simpler setup and takes advantageof recording the large initial ultrasound pressures without appreciableacoustic loses.

During in vivo imaging experiments, no agent or ultrasound couplingmedium are required. However the target can be prepared with water orany liquid such as oil before non-contact imaging session. NI-PARS doesnot require a floating table unlike many other interferometric sensors.No special holder or immobilization is required to hold the targetduring imaging sessions.

Other advantages that are inherent to the structure will be apparent tothose skilled in the art. The embodiments described herein areillustrative and not intended to limit the scope of the claims, whichare to be interpreted in light of the specification as a whole.

A pulse laser is used to generate photoacoustic signals and the acousticsignatures are interrogated using either a long-coherence orshort-coherence length probe beam co-focused with the excitation spots.The NI-PARS system is utilized to remotely record the large localinitial pressures from chromophores and without appreciable acousticloses due to diffraction, propagation and attenuation.

The excitation beam may be any pulsed or modulated source ofelectromagnetic radiation including lasers or other optical sources. Inone example, a nanosecond-pulsed laser was used. The excitation beam maybe set to any wavelength suitable for taking advantage of optical (orother electromagnetic) absorption of the sample. The source may bemonochromatic or polychromatic.

The interrogation beam may be any pulsed, continues or modulated sourceof electromagnetic radiation including lasers or other optical sources.Any wavelength can be sued for interrogation purpose depends on theapplication.

The chromatic aberration in the collimating and objective lens pair washarnessed to refocus light from a fiber into the object so that eachwavelength is focused at a slightly different depth location. Usingthese wavelengths simultaneously was previously shown to improve thedepth of field and SNR for structural imaging of microvasculature withOR-PAM.

Since the design is not interferometric, theprobe/receiver/interrogation beam, may be a long-coherence or ashort-coherence length probe beam. Without need of any reference beam orreference arm. Using a short-coherence length, however, may ensurepreclusion of interference from reflections in the system or sample toavoid unwanted signals and to extract only photoacoustic initialpressures.

Unlike optical coherence tomography (OCT) or interferometry detection ofphotoacoustic signal, the NI-PARS system detects the changes in theamount of the reflected light from sample due to acoustic pressure andno interferometry design such as, reference beam, reference arm or axialscanning of reference beam are needed.

NI-PARS may be integrated with OCT to provide a complete set ofinformation offered by both photoacoustic and OCT systems.

NI-PARS with a short or long-coherence beam may be used for eitheroptical resolution photoacoustic microscopy (OR-PAM) or commonphotoacoustic microscopy (PAM).

In one example, both excitation and receiver beam may be combined andscanned. In this way, photoacoustic excitations may be sensed in thesame area as they are generated and where they are the largest. Otherarrangements may also be used, including keeping the receiver beam fixedwhile scanning the excitation beam or vice versa. Galvanometers, MEMSmirrors and stepper/DC motors may be used as a means of scanning theexcitation beam, probe/receiver beam or both.

The configurations shown in FIGS. 4 a - 4 d may be used to performNI-PARS imaging. In the depicted embodiments, lines 502, depicted with alarger radius of curvature, represent excitation beams and lines 504,depicted with a smaller radius of curvature, represent receiver beams.FIG. 5 a offers a kind of confocal photoacoustic system where theexcitation beam 502 and probing receive beam 504 are focused on the samespot, which can be on a micron- or submicron scale. In FIG. 5 b , theoptical resolution can be provided by the receiver beam 504, rather thanthe excitation beam 502. FIG. 5 c shows excitation beam 502 and receiverbeam 502 focused on different spots, and takes advantage of ultrasoundtime of flight in order to locate the excitation and receiver beams 502and 504 at different positions. In FIG. 4 d , optical resolution isprovided by the excitation beam 502. Preferably, the focus of either orboth of the excitation beam 502 or the detection beam 504 is less than30 µm, less than 10 µm, less than 1 µm, or to the diffraction limit oflight. A tighter focus results in a higher possible resolution and abetter signal to noise ratio in the reflected beam that is detected. Asused herein, the term “focus” is intended to refer to the focal zone ofthe beam, or the point at which the beam spot size is at the tightestsize, and where the diameter of the focal zone is 30% greater than thediameter of the beam spot size. Also preferably, the excitation anddetection beams 502 and 504 are focused on the same position, althoughthere may be some spacing between the respective focuses as shown inFIG. 4 c . In FIG. 4 c , the beams may be focused at differentlocations, but preferably within 1 mm, 0.5 mm, 100 µm or within therange of the largest focus of the beam. In FIGS. 4 a, 4 b and 4 d , thebeams may be confocal, or may overlap within the focus of the beam withthe largest focus. For example, in FIG. 4 a , the excitation beam islarger than the detection beam, and the detection beam is directed at alocation within the focus of the excitation beam. By moving thedetection beam, the area within the excitation beam may be imaged. Byhaving confocal beams, both beams may be moved to image the sample.

One or both of the beams are preferably focused below the surface of thesample. Generally speaking, the beams may be effectively focused up to 1mm below the surface of the sample. The beams may be focused at least 1µm below the surface, or focused such that focal point of the beam is atleast the distance of focal zone of the beam below the surface of thesample. It will be understood that, while both beams are preferablyfocused below the surface, in some embodiments either the excitationbeam or the interrogation beam may be focused below the surface, withthe other focused on, for example, the surface of the sample. In caseswhere only one beam is focused below the surface of the sample, theseparation between the beams discussed previously will be a lateralseparation, i.e. in the plane of the sample and orthogonal to the depthof the sample.

The excitation beam and sensing/receiver beam can be combined usingdichroic mirrors, prisms, beamsplitters, polarizing beamsplitters etc.They can also be focused using different optical paths.

The reflected light may be collected by photodiodes, avalanchephotodiodes, phototubes, photomultipliers, CMOS cameras, CCD cameras(including EM-CCD, intensified-CCDs, back-thinned and cooled CCDs), etc.The detected light may be amplified by an RF amplifier, lock-inamplifier, trans-impedance amplifier, or other amplifier configuration.Also different methods may be used in order to filter the excitationbeam from the receiver beam before detection. NI-PARS may use opticalamplifiers to amplify detected light.

NI-PARS can be used in many form factors, such as table top, handheldand endoscopy. Examples of endoscopy NI-PARS are shown in FIGS. 5 a, 5 band 5 c with various arrangements of NI-PARS excitation units 1102,NI-PARS detection units 1104, fibre optics 1106 such as image-guidefibers, and lenses 1108 that focus the respective beams onto the sample18. When excitation and detection units 1102 and 1104 are separated,there may be a separate fibre 1110 provided, such as a single modefiber.

A table top and handheld NI-PARS system may be constructed based onprinciples known in the art. The proposed NI-PARS system takes advantageof optical excitation and detection which can help to dramaticallyreduce the footprint of the system. The footprint of previous systemshas been much too large to use the system in all but body surfaces. Forendoscopic applications, the footprint of the ultrasound detector mustbe minimized to make the imaging catheter small and flexible enough tonavigate through small orifices and vessels. The piezoelectric receiversare not ideal candidates for endoscopic applications as there istrade-off between the sensitivity and the size of the receiver. On theother hand for many invasive applications sterilizable or disposablecatheters and a non-contact approach are necessary. The system may alsobe used as NI-PARS endoscopy system with a potential footprint the sizeof an optical fiber, as both excitation and NI-PARS beam can be coupledinto a single mode fiber or image guide fiber.

Image-guide fibers (miniaturized fiber bundles with as many as 100,000or more individual micrometer-sized strands in a single optical fiberwith diameters ranging from 200 µm to 2 mm) may be used to transmit bothfocused light spots. The excitation beam may be scanned either at thedistal end or proximal end of the fiber using one of the scanningmethods mentioned before. However, the receiver beam may be scanned orbe fixed. The scanned spot is transmitted via the image-guide fiber 1106to the output end. Therefore, it may be used to directly contact thesample, or re-focused using an attached miniature GRIN lens 1108. In oneexample, C-scan photoacoustic images were obtained from the fiberimage-guides using an external ultrasound transducer to collectphotoacoustic signals. Using an edge-spread and Gaussian function, aresolution of approximately 7 µm was obtained using the image-guidefiber 1106. It is believed that a higher resolution may also be obtainedwith appropriate improvements to the setup and equipment used.

The NI-PARS system may be combined with other imaging modalities such asfluorescence microscopy, two-photon and confocal fluorescencemicroscopy, Coherent-Anti-Raman-Stokes microscopy, Raman microscopy,Optical coherence tomography, other photoacoustic and ultrasoundsystems, etc. This could permit imaging of the microcirculation, bloodoxygenation parameter imaging, and imaging of other molecularly-specifictargets simultaneously, a potentially important task that is difficultto implement with only fluorescence based microscopy methods. An exampleof a NI-PARS system 10 integrated with another optical imaging system1202 is shown in FIG. 6 , where NI-PARS 10 and the other optical imagingsystem 1202 are both connected to the sample 18 by a combiner 1204.

NI-PARS can be integrated with any interferometric designs for detectionphotoacoustic signals to extent its application. Interferometric designssuch as common path interferometer (using specially designedinterferometer objective lenses), Michelson interferometer, Fizeauinterferometer, Ramsey interferometer, Sagnac interferometer,Fabry-Perot interferometer and Mach-Zehnder interferometer.

NI-PARS may be used for A, B or C scan images for in vivo, ex vivo orphantom studies.

A multi-wavelength visible laser source may also been implemented togenerate photoacoustic signals for functional or structural imaging.

NI-PARS may be optimized in order to takes advantage of a multi-focusdesign for improving the depth-of-focus of 2D and 3D OR-NI-PARS imaging.The chromatic aberration in the collimating and objective lens pair maybe harnessed to refocus light from a fiber into the object so that eachwavelength is focused at a slightly different depth location. Usingthese wavelengths simultaneously may be used to improve the depth offield and signal to noise ratio (SNR) of NI-PARS images. During NI-PARSimaging, depth scanning by wavelength tuning may be performed.

Polarization analyzers may be used to decompose detected light intorespective polarization states. The light detected in each polarizationstate may provide information about ultrasound-tissue interaction.

Applications

It will be understood that the system described herein may be used invarious ways, such as those purposes described in the prior art, andalso may be used in other ways to take advantage of the aspectsdescribed above. A non-exhaustive list of applications is discussedbelow.

The system may be used for imaging angiogenesis for differentpre-clinical tumor models.

The system may also be used for clinical imaging of micro- andmacro-circulation and pigmented cells, which may find use forapplications such as in (1) the eye, potentially augmenting or replacingfluorescein angiography; (2) imaging dermatological lesions includingmelanoma, basal cell carcinoma, hemangioma, psoriasis, eczema,dermatitis, imaging Mohs surgery, imaging to verify tumor marginresections; (3) peripheral vascular disease; (4) diabetic and pressureulcers; (5) burn imaging; (6) plastic surgery and microsurgery; (7)imaging of circulating tumor cells, especially melanoma cells; (8)imaging lymph node angiogenesis; (9) imaging response to photodynamictherapies including those with vascular ablative mechanisms; (10)imaging response to chemotherapeutics including anti-angiogenic drugs;(11) imaging response to radiotherapy.

The system may be useful in estimating oxygen saturation usingmulti-wavelength photoacoustic excitation and NI-PARS detection andapplications including: (1) estimating venous oxygen saturation wherepulse oximetry cannot be used including estimating cerebrovenous oxygensaturation and central venous oxygen saturation. This could potentiallyreplace catheterization procedures which can be risky, especially insmall children and infants.

Oxygen flux and oxygen consumption may also be estimated by usingNI-PARS imaging to estimate oxygen saturation, and an auxiliary methodto estimate blood flow in vessels flowing into and out of a region oftissue.

The system may also have some gastroenterological applications, such asimaging vascular beds and depth of invasion in Barrett’s esophagus andcolorectal cancers. Depth of invasion is key to prognosis and metabolicpotential. Gastroenterological applications may be combined orpiggy-backed off of a clinical endoscope and the miniaturized NI-PARSsystem may be designed either as a standalone endoscope or fit withinthe accessory channel of a clinical endoscope.

The system may have some surgical applications, such as functionalimaging during brain surgery, use for assessment of internal bleedingand cauterization verification, imaging perfusion sufficiency of organsand organ transplants, imaging angiogenesis around islet transplants,imaging of skin-grafts, imaging of tissue scaffolds and biomaterials toevaluate vascularization and immune rejection, imaging to aidmicrosurgery, guidance to avoid cutting critical blood vessels andnerves.

Other examples of applications may include NI-PARS imaging of contrastagents in clinical or pre-clinical applications; identification ofsentinel lymph nodes; non- or minimally-invasive identification oftumors in lymph nodes; imaging of genetically-encoded reporters such astyrosinase, chromoproteins, fluorescent proteins for pre-clinical orclinical molecular imaging applications; imaging actively or passivelytargeted optically absorbing nanoparticles for molecular imaging; andimaging of blood clots and potentially staging the age of the clots.

NI-PARS Mechanism

Rather than calculate the phase shifts of transmitted light NI-PARS isinterested in the light reflected from a refractive index mismatch. Witha large initial pressure a significate refractive index step changes inthe confined excitation volume occurs. This results in a large amplitudereflection coefficient. This mechanism will contribute to both amplitudeand phase variations in the probe beam.

Refractive index changes from their unperturbed state can occur due topressure rises. Initial pressures generated by the absorption of anoptical excitation pulse which is shorter in time than the stressconfinement and thermal confinement condition is described by p₀ = Γϕµawhere Γ is a material property known as the Grüneissen parameter, ϕ isthe focal fluence of the excitation beam and µa is the opticalabsorption of the medium at the given excitation wavelength. Thesepressures can be large. As an example, assuming a focal fluence of 500mJcm⁻ ², a whole blood sample (with a hemoglobin concentration of 150gL⁻¹ assuming a fully oxygenated sample) with optical absorption at 532nm excitation of 237 cm⁻¹ and a Grüneissen parameter of 1 gives aninitial pressure of 119 MPa. These pressures are sufficient to create ameasureable change in the refractive index which follows theelasto-optic relation such that the perturbed refractive index (n*(r,t)) is related to the unperturbed refractive index (n(r, t)) and thepressure field (p(r, t)) by

n * (r, t) = n[1 + (ηn²p/2pv_(s)²)]

where η is the elasto-optic coefficient, ρ is the specific density andν_(s) is the speed of sound.

The accumulated phase shift of light passing through a zone of enhancedpressure can be calculated by Raman Nath diffraction theory and willdepend on the direction of the sound and the direction of the light aswell as the pressure field inhomogeneity. For a light beam incident on aplane pressure wave where both the light and sound beams are parallel,the accumulated phase shift should be zero and are rather maximum whensound fields create effective diffraction gratings orthogonal to thelight propagation. The electric field back-reflected from the sample andincident on the photodiode is modelled as having two components, AC andDC terms ES=E(DC)+E(AC) Here E(DC) is the electric field magnitude oflight reflected from the sample surface and E(AC) is the electric fieldamplitude of light reflected from the excitation volume beneath thesurface due to a transient pressure induced optical index step. Thefraction of light modulated compared to surface reflected light FP iscalculated as FP=(IE(AC) |^2 〉/〈|E(DC) |^2 ). The fraction ofmodulated light due to pressure-induced refractive index change can beas significant of several percentage of the incident light.

When the sample consists of a planar interface (where the curvature ofthe surface is much larger than the probe beam focal spot size) thesmall changes in refractive index can produce changes in thereflectivity of the probe beam. This can be modeled by taking themdifference in the intensity reflection coefficients of the perturbed

$\left( {R^{*} = \left| \frac{n_{1 + \delta n - n_{2}}}{n_{1 + \delta n + n_{2}}} \right|^{2}} \right)$

and unperturbed

$\left( {R_{s} = \left| \frac{n_{1 - n_{2}}}{n_{1 + n_{2}}} \right|^{2}} \right)$

interfaces such that the measured signal varies with amplitude followingΔR = R* - R_(s) where n₁ is the refractive index of the opticalabsorbing medium and n₂ is the refractive index of the opticallytransparent medium. Assuming that the perturbation is small such that

δn ≪ 1; δn ≪ n_(1,)n₂

and that the refractive indices are all mostly real (this is not arequirement, but merely allows for intuitive approximation) provides anexpression following

$\Delta R = 4\delta nR_{s}\frac{n_{2}}{n_{1}^{2} - n_{2}^{2}} + \mathcal{O}\left\{ {\delta n} \right\}$

where

𝒪{δn}

represents higher order terms. This describes a linear relationship withthe static reflectivity (due to the initial refractive index mismatchbetween n₁ and n₂) and a linear relationship with the perturbation

δn

which is in turn proportional to the photoacoustic initial pressure.

In the case of an object, or a collection of objects which are opticallyscattering, optically absorbing, and are on the scale of or smaller thanthe beam spot size, optical property changes are more appropriatelydescribed by scattering theory. The refractive index modulation broughton by the photoacoustic initial pressures now is assumed to alter thescattering properties of individual particles. For the case of singleparticle interactions this can be observed as a change in the collectedfraction (or observed reflection) such that

ΔR_(c) = R_(c)^(*) − R_(s, c)

where

R_(s)^(*)

is the modified collected fraction and

R_(s, c)

is the unperturbed collected fraction. In brief these describe arelationship which follows

ΔR_(c) ∝ Δσ_(s)

where

$\Delta\sigma_{s} = 2\delta n_{2}\sigma_{s}\left( {\frac{n_{s} - n_{b}}{\left| {n_{s} - n_{b}} \right|^{2}} - 1} \right) + \mathcal{O}\left\{ {\delta n} \right\}$

is the change in the scattering cross-section of the particle,

n_(s)

is the unperturbed refractive index of the particle,

δn_(s)

is the refractive index change brought on by the elasto-optic effect,

n_(b)

is the refractive index of the background medium and

σ_(s)

is the unperturbed scattering cross section. In the case of a largeensemble of particles, the change in diffuse reflection is insteadmonitored which is described by radiative transfer theory in whichindividual particles are assumed to form an equivalent homogenousscattering medium with a perturbed and unperturbed set of opticalscattering properties.

The system is sensitive to intensity reflectivity modulations at anydepth within the probe beam optical depth-of-focus. Such modulationseffectively begin instantaneously, coincident with the excitation pulse,irrespective of depth. Because the proposed system reads outphase-insensitive intensity reflectivity, time-resolved signals do notproduce depth-resolved information.

Experimental Results

FIG. 7 a shows NI-PARS (using experimental setup shown in FIG. 2 )imaging of carbon fiber networks using ~1 nJ excitation pulse energy and4 mW interrogation power. SNR (defined as average of signal over thestandard deviation of the noise) was quantified as 45 dB. FIG. 7 b showsFWHM due to fitting individual carbon fiber (with ~6 µm diameter) signalamplitude to a Gaussian function. FIG. 7 c shows the resolution studyusing a knife edge spread function. The lateral resolution of the systemhas been measured as ∼2.5±1 µm.

FIG. 8 show in vivo images of CAM-membrane of 5-day chicken embryosusing the experimental setup shown in FIG. 1 FIG. 8 shows multi focusNI-PARS images revealing both capillary beds and bigger blood vessels.In the chicken embryo model bigger blood vessels usually are locateddeeper than capillaries. In order to see both deep- and shallow vesselssimultaneously the multi-focus design is optimized to extend thedepth-of-field to ∼ 250 µm. A single wavelength can be used as well asshown in FIG. 7 , However, with a single wavelength, depth-of-focus islimited to ∼30 µm, rather than 250 µm with the multi-focus approach.Hence single-wavelength excitation is better-suited fordepth-sectioning. FIG. 10 depicts the NI-PARS frequency response andFIG. 9 depicts the NI-PARS photoacoustic time domain signal of anindividual carbon fiber.

FIG. 11 (using experimental setup shown in FIG. 1 and FIG. 2 ) depict invivo NI-PARS images of a mouse ear. In all in vivo images pulse energy∼20-80 nJ was used and the interrogation power was fixed to 6 mW.

All the images shown herein are raw data and no major image processingsteps are applied.

As will be understood, the high sensitivity and the fine resolution ofthe proposed system offer performance comparable to other in vivooptical resolution photoacoustic microscopy systems but with much highersignal to noise ratio and in a non-contact reflection mode suitable formany clinical and pre-clinical applications.

In this patent document, the word “comprising” is used in itsnon-limiting sense to mean that items following the word are included,but items not specifically mentioned are not excluded. A reference to anelement by the indefinite article “a” does not exclude the possibilitythat more than one of the elements is present, unless the contextclearly requires that there be one and only one of the elements.

The scope of the following claims should not be limited by the preferredembodiments set forth in the examples above and in the drawings, butshould be given the broadest interpretation consistent with thedescription as a whole.

1-16. (canceled)
 17. A method of imaging a sample, comprising: receiving, at one or more processors, information relating to a non-interferometrically detected portion of an interrogation beam, wherein the detected portion of the interrogation beam was detected using a non-interferometric detector, wherein the detected portion of the beam returned from a sample interrogated with the interrogation beam and excited with an excitation beam to generate signals in the sample at an excitation location, wherein the interrogation beam was directed to the sample at or adjacent to the excitation location, and wherein at least one of the excitation beam or the interrogation beam were focused below a surface of the sample; and generating or calculating, by the one or more processors, an image of the sample based on the received information.
 18. The method of claim 17, wherein generating or calculating the image is based on a detected intensity modulation of the detected portion of the interrogation beam, wherein the non-interferometric detector is configured to preclude phase-modulation sensitivity to enable detection of intensity variations.
 19. The method of claim 17, wherein the interrogation beam or the excitation beam are focused within 1 mm of the surface of the sample.
 20. The method of claim 17, wherein at least one of the interrogation beam or the excitation beam is focused at a depth greater than 1 µm below the surface of the sample.
 21. The method of claim 17, wherein the interrogation beam is focused at a first focal point or the excitation beam is focused at a second focal point, the first or second focal points being below the surface of the sample.
 22. The method of claim 21, wherein at least one of the first or second focal points are spaced below the surface of the sample at a depth that is greater than a focal zone of a respective at least one of the interrogation beam or the excitation beam.
 23. The method of claim 17, wherein the interrogation beam and the excitation beam have a separation of less than 1 mm within the sample.
 24. The method of claim 17, wherein the excitation beam has a focal point that is within a focal zone of the interrogation beam; or the interrogation beam has a focal point that is within a focal zone of the excitation beam.
 25. The method of claim 17, wherein the excitation beam is scanned through the sample while the interrogation beam is stationary.
 26. The method of claim 17, wherein the interrogation beam is scanned through the sample while the excitation beam is stationary.
 27. The method of claim 17, wherein both the interrogation beam and the excitation beam are scanned through the sample concurrently.
 28. The method of claim 17, wherein at least one of the interrogation beam or the excitation beam has a focal diameter of less than 30 µm.
 29. The method of claim 17, wherein the method is used for estimating blood flow in vessels flowing into and out of a region of tissue.
 30. The method of claim 17, wherein the method is used for estimating oxygen saturation in the sample.
 31. The method of claim 17, wherein the method is used in one or more of the following applications: imaging angiogenesis for pre-clinical tumor models; estimating oxygen saturation using multi-wavelength photoacoustic excitation; estimating venous oxygen saturation where pulse oximetry cannot be used; estimating cerebrovenous oxygen saturation and/or central venous oxygen saturation; estimating oxygen flux and/or oxygen consumption; clinical imaging of micro- and macro-circulation and pigmented cells; imaging of the eye; augmenting or replacing fluorescein angiography; imaging dermatological lesions; imaging melanoma; imaging basal cell carcinoma; imaging hemangioma; imaging psoriasis; imaging eczema; imaging dermatitis; imaging Mohs surgery; imaging to verify tumor margin resections; imaging peripheral vascular disease; imaging diabetic and/or pressure ulcers burn imaging; plastic surgery; microsurgery; imaging of circulating tumor cells; imaging melanoma cells; imaging lymph node angiogenesis; imaging response to photodynamic therapies; imaging response to photodynamic therapies having vascular ablative mechanisms; imaging response to chemotherapeutics; imaging response to anti-angiogenic drugs; imaging response to radiotherapy; imaging histology; imaging pathology specimen; imaging vascular beds and depth of invasion in Barrett’s esophagus and/or colorectal cancers; functional imaging during brain surgery; assessment of internal bleeding and/or cauterization verification; imaging perfusion sufficiency of organs and/or organ transplants; imaging angiogenesis around islet transplants; imaging of skin-grafts; imaging of tissue scaffolds and/or biomaterials to evaluate vascularization and/or immune rejection; imaging to aid microsurgery; guidance to avoid cutting blood vessels and/or nerves; imaging of contrast agents in clinical or pre-clinical applications; identification of sentinel lymph nodes; non- or minimally-invasive identification of tumors in lymph nodes; imaging of genetically-encoded reporters, wherein the genetically-encoded reporters include tyrosinase, chromoproteins, and/or fluorescent proteins for pre-clinical or clinical molecular imaging applications; imaging actively or passively targeted optically absorbing nanoparticles for molecular imaging; imaging of blood clots; or staging an age of blood clots.
 32. The method of claim 17, wherein the non-interferometric detector is configured to sense a pressure-induced refractive-index modulation or a temperature-induced refractive-index modulation.
 33. The method of claim 17, wherein the non-interferometric detector is not sensitive to scattered probe beam phase modulations associated with motion of scatterers, and is not sensitive to subsurface and surface oscillations.
 34. The method of claim 17, further comprising amplifying an existing refractive index where absorption is present to detect change in intensity reflectivity.
 35. A non-interferometric photoacoustic remote sensing system (NI-PARS) for imaging a subsurface structure in a sample, comprising: a processor configured to calculate an image of a sample based on a detected intensity modulation of a returning portion of an interrogation beam from below a surface of the sample, wherein: the sample was excited using an excitation beam configured to generate signals in the sample at an excitation location; the sample was interrogated using the interrogation beam incident on the sample at the excitation location, wherein the returning portion of the interrogation beam is indicative of the generated signals; the excitation beam and the interrogation beam were focused below a surface of the sample; and the returning portion of the interrogation beam was detected using a non-interferometric detector configured for non-interferometric .
 36. A method of imaging a sample, comprising: calculating an image of a sample based on a detected intensity modulation of a returning portion of an interrogation beam from below a surface of the sample, wherein: the returning portion of the interrogation beam was non-interferometrically detected using a using a non-interferometric detector configured for non-interferometric sensing, the sample was excited using an excitation beam configured to generate signals in the sample at an excitation location; the sample was interrogated using the interrogation beam, wherein the returning portion of the interrogation beam is indicative of the generated signals; and at least one of the excitation beam or the interrogation beam were focused below a surface of the sample. 